Endoscope

ABSTRACT

Disclosed herein is an endoscope comprising: an insertion tube; a radiation detector configured to detect radiation particles in a first range of energy and radiation particles in a second range of energy.

TECHNICAL FIELD

The disclosure herein relates to an endoscope, particularly relates to an endoscope comprising a radiation detector.

BACKGROUND

An endoscope is a tubular medical instrument used to examine inside a human body or internal organs by inserting through a small cut, or an opening in the body such as the mouth. It may be used to examine, diagnosis, or assist in the surgery. With attached camera and other accessories like forceps and scissors, the endoscope may allow surgeons to view or operate on the internal organs and blood vessels of human body. Examples of endoscopes include the cystoscope, nephroscope, bronchoscope, arthroscope, colonoscope, and laparoscope, etc.

Radiation fluorescence is the emission of characteristic fluorescent radiation from a material that has been excited by, for example, exposure to high-energy X-rays or gamma rays. An electron on an inner orbital of an atom may be ejected, leaving a vacancy on the inner orbital, if the atom is exposed to X-rays or gamma rays with photon energy greater than the ionization potential of the electron. When an electron on an outer orbital of the atom relaxes to fill the vacancy on the inner orbital, a radiation is emitted. The emitted radiation has a photon energy equal the energy difference between the outer orbital and inner orbital electrons.

For a given atom, the number of possible relaxations is limited. As shown in FIG. 1A, when an electron on the L orbital relaxes to fill a vacancy on the K orbital (L→K), the fluorescent radiation is called Kα. The fluorescent radiation from M→K relaxation is called Kβ. As shown in FIG. 1B, the fluorescent radiation from M→L relaxation is called Lα, and so on.

Semiconductor radiation detectors can directly convert radiation into electric signals. A semiconductor radiation detector may include a semiconductor layer that absorbs radiation in wavelengths of interest. When a radiation particle is absorbed in the semiconductor layer, multiple charge carriers (e.g., electrons and holes) are generated whose amount is proportional to the energy of the radiation. As used herein, the term “charge carriers,” “charges” and “carriers” are used interchangeably. The charge carriers are collected and counted by a circuitry to determine the energy of the radiation and the process repeats for the next incident radiation. A spectrum may be compiled by counting the number of detected radiation as a function of its energy.

Analyzing the fluorescent radiation spectrum can identify the elements in a sample because each element has orbitals of characteristic energy. The intensity of each characteristic energy peak is directly related to the amount of each element in the sample.

SUMMARY

Disclosed herein is an endoscope comprising: an insertion tube; a radiation detector configured to detect radiation particles in a first range of energy and radiation particles in a second range of energy.

According to an embodiment, the radiation particles are X-ray photons.

According to an embodiment, the radiation particles are emitted by fluorescence of a tissue in a person.

According to an embodiment, the tissue is a blood vessel.

According to an embodiment, the radiation detector is configured to count numbers of radiation particles incident thereon whose energy fall in a plurality of bins, within a period of time.

According to an embodiment, the radiation detector is configured to determine existence of a chemical element based on intensity of the radiation particles in the first range of energy and intensity of the radiation particles in the second range of energy.

According to an embodiment, the chemical element is calcium.

According to an embodiment, the radiation detector is configured to determine a spectrum of the radiation particles incident on the radiation detector.

According to an embodiment, the radiation detector comprises a plurality of pixels, wherein each of the pixels is configured to count numbers of radiation particles incident thereon whose energy fall in a plurality of bins, within a period of time; and wherein the radiation detector is configured to determine a sum of the numbers of radiation particles for the bins of the same energy range counted by all the pixels.

According to an embodiment, the pixels are configured to count the numbers of X-ray photons within a same period of time.

According to an embodiment, each of the pixels comprises an analog-to-digital converter (ADC) configured to digitize an analog signal representing the energy of an incident radiation particle into a digital signal.

According to an embodiment, the pixels are configured to operate in parallel.

According to an embodiment, the ADC is a successive-approximation-register (SAR) ADC.

According to an embodiment, the radiation detector comprises a radiation absorption layer comprising an electric contact; a first voltage comparator configured to compare a voltage of the electric contact to a first threshold; a second voltage comparator configured to compare the voltage to a second threshold; a plurality of counters each associated with a bin and configured to register a number of radiation particles absorbed by the radiation absorption layer wherein the energy of the radiation particles falls in the bin; a controller; wherein the controller is configured to start a time delay from a time at which the first voltage comparator determines that an absolute value of the voltage equals or exceeds an absolute value of the first threshold; wherein the controller is configured to determine whether an energy of a radiation particle falls into the bin; wherein the controller is configured to cause the number registered by the counter associated with the bin to increase by one.

According to an embodiment, the endoscope further comprises a capacitor module electrically connected to the electric contact, wherein the capacitor module is configured to collect charge carriers from the electric contact.

According to an embodiment, the controller is configured to activate the second voltage comparator at a beginning or expiration of the time delay.

According to an embodiment, the controller is configured to connect the electric contact to an electrical ground.

According to an embodiment, the rate of change of the voltage is substantially zero at expiration of the time delay.

According to an embodiment, the radiation absorption layer comprises a diode.

According to an embodiment, the radiation absorption layer comprises silicon or germanium, GaAs, CdTe, CdZnTe, or a combination thereof.

According to an embodiment, the radiation detector does not comprise a scintillator.

Disclosed herein is a system comprising the endoscope described herein and a radiation source.

Disclosed herein is a method comprising: inserting the endoscope described herein into a human body tissue, for example, a blood vessel; detecting, using the radiation detector, an intensity of radiation in a first range of energy and an intensity of radiation in a second range of energy; determining existence of a chemical element in the human body based on the intensity of radiation in the first range of energy and the intensity of radiation in the second range of energy.

BRIEF DESCRIPTION OF FIGURES

FIG. 1A and FIG. 1B schematically show mechanisms of radiation fluorescence.

FIG. 2 schematically shows a cross-sectional view of an endoscope with a radiation detector enclosed, according to an embodiment.

FIG. 3 schematically shows a top view of the radiation detector and its plurality of pixels, according to an embodiment.

FIG. 4A schematically shows a block diagram for the radiation detector, according to an embodiment.

FIG. 4B schematically shows a detailed cross-sectional view of the radiation detector.

FIG. 4C schematically shows an alternative detailed cross-sectional view of the radiation detector.

FIG. 5A and FIG. 5B each schematically show a component diagram of the electronic system of the radiation detector, according to an embodiment.

FIG. 6 schematically shows a temporal change of the electric current flowing through an electric contact (upper curve) caused by charge carriers generated by a radiation photon incident on a pixel associated with the electric contact, and a corresponding temporal change of the voltage of the electric contact (lower curve).

FIG. 7 shows an example flow chart for an application of the endoscope with the radiation detector, according to an embodiment.

FIG. 8 schematically shows a human tissue, i.e. blood vessel, scanning endoscope comprising a radiation detector described herein.

DETAILED DESCRIPTION

FIG. 2 schematically shows a cross-sectional view of an endoscope 101 with a radiation detector 100 enclosed, according to an embodiment. The endoscope 101 may comprise an insertion tube 102, which may be rigid or flexible, a signal cable 103, and a controlling unit 104 which may be configured to receive or transmit signals, or control the movement of the insertion tube 102. The insertion tube 102 may have a very small diameter (e.g., less than 1 mm), which is suitable for inserting into a human body or tissues, like blood vessels. The insertion tube 102 may be made of materials that is transparent to the radiation of interest, and may encapsulate the radiation detector 100 inside. The radiation detector 100 may include a radiation absorption layer 110 and an electronics layer 120 (e.g., an ASIC) for processing or analyzing electrical signals incident radiation generates in the radiation absorption layer 110. The radiation absorption layer 110 may include a semiconductor material such as, silicon, germanium, GaAs, CdTe, CdZnTe, or a combination thereof. The semiconductor may have a high mass attenuation coefficient for the radiation energy of interest.

According to an embodiment, the radiation detector 100 of the endoscope 101 is configured to collect radiation particles of radiation fluorescence from a human tissue while the endoscope 101 is inside a human body. The radiation particles may be X-ray photons. The endoscope 101 may be inserted into a human blood vessel and used to determine the existence of a chemical element (e.g., Ca) by analyzing an X-ray photon energy spectrum collected by the radiation detector 100. For example, the endoscope may be used to find local calcification by determining the presence and concentration of calcium. The radiation detector 100 is configured to detect and count the X-ray photons emitted from calcium due to radiation fluorescence in a first range of energy, i.e. the Kα photon energy peak, and detect and count the X-ray photons emitted from Calcium in a second range of energy, i.e. the Kβ photon energy peak.

FIG. 3 schematically shows a top view of the radiation detector 100, according to an embodiment. The radiation detector 100 may have an array of pixels 150. The array may be a rectangular array, a honeycomb array, a hexagonal array or any other suitable array. Each pixel 150 is configured to detect a radiation photon incident thereon and measure the energy of the radiation particle. For example, each pixel 150 is configured to count numbers of radiation particles incident thereon whose energy falls in a plurality of bins, within a period of time. All the pixels 150 may be configured to count the numbers of radiation particles incident thereon within a plurality of bins of energy within the same period of time. Each pixel 150 may have its own analog-to-digital converter (ADC) configured to digitize an analog signal representing the energy of an incident radiation particle into a digital signal. For XRF applications, an ADC with a 10-bit resolution or higher is useful. Each pixel 150 may be configured to measure its dark current, such as before or concurrently with each radiation particle incident thereon. Each pixel 150 may be configured to deduct the contribution of the dark current from the energy of the radiation particle incident thereon. The pixels 150 may be configured to operate in parallel. For example, when one pixel 150 measures an incident radiation particle, another pixel 150 may be waiting for a radiation particle to arrive. The pixels 150 may not have to be individually addressable.

Due to the small size of the insertion tube 102 of the endoscope 101, the radiation detector 100 may have a single or an array of a few pixels 150 (for example, 2 or 4). The radiation detector 100 may be configured to add the numbers of radiation particles for the bins of the same energy range counted by all the pixels 150. For example, the radiation detector 100 may add the numbers the pixels 150 stored in a bin for energy from 70 KeV to 71 KeV, add the numbers the pixels 150 stored in a bin for energy from 71 KeV to 72 KeV, and so on. The radiation detector 100 may compile the added numbers for the bins as a spectrum of the radiation particles incident on the radiation detector 100.

FIG. 4A schematically shows a block diagram for the radiation detector 100, according to an embodiment. Each pixel 150 may measure the energy 151 of a radiation particle incident thereon. The energy 151 of the radiation particle is digitized (e.g., by an ADC) in step 152 into one of a plurality of bins 153A, 153B, 153C . . . . The bins 153A, 153B, 153C . . . each have a corresponding counter 154A, 154B and 154C, respectively. When the energy 151 is allocated into a bin, the number stored in the corresponding counter increases by one. The radiation detector 100 may added the numbers stored in all the counters corresponding to bins for the same energy range in the pixels 150. For example, the numbers stored in all the counters 154C in all pixels 150 may be added and stored in a global counter 100C for the same energy range. The numbers stored in all the global counters may be compiled into an energy spectrum of the radiation incident on the radiation detector 100.

As shown in a detailed cross-sectional view of the radiation detector 100 in FIG. 4B, according to an embodiment, the radiation absorption layer 110 may include one or more diodes (e.g., p-i-n or p-n) formed by a first doped region 111, one or more discrete regions 114 of a second doped region 113. The second doped region 113 may be separated from the first doped region 111 by an optional the intrinsic region 112. The discrete regions 114 are separated from one another by the first doped region 111 or the intrinsic region 112. The first doped region 111 and the second doped region 113 have opposite types of doping (e.g., region 111 is p-type and region 113 is n-type, or region 111 is n-type and region 113 is p-type). In the example in FIG. 4B, each of the discrete regions 114 of the second doped region 113 forms a diode with the first doped region 111 and the optional intrinsic region 112. Namely, in the example in FIG. 4B, the radiation absorption layer 110 has a plurality of diodes having the first doped region 111 as a shared electrode. The first doped region 111 may also have discrete portions.

When a radiation particle hits the radiation absorption layer 110 including diodes, the radiation particle may be absorbed and generate one or more charge carriers by a number of mechanisms. A radiation particle may generate 10 to 100000 charge carriers. The charge carriers may drift to the electrodes of one of the diodes under an electric field. The field may be an external electric field. The electrical contact 119B may include discrete portions each of which is in electrical contact with the discrete regions 114. In an embodiment, the charge carriers may drift in directions such that the charge carriers generated by a single radiation particle are not substantially shared by two different discrete regions 114 (“not substantially shared” here means less than 2%, less than 0.5%, less than 0.1%, or less than 0.01% of these charge carriers flow to a different one of the discrete regions 114 than the rest of the charge carriers). Charge carriers generated by a radiation particle incident around the footprint of one of these discrete regions 114 are not substantially shared with another of these discrete regions 114. A pixel 150 associated with a discrete region 114 may be an area around the discrete region 114 in which substantially all (more than 98%, more than 99.5%, more than 99.9%, or more than 99.99% of) charge carriers generated by an X-ray photon incident therein flow to the discrete region 114. Namely, less than 2%, less than 1%, less than 0.1%, or less than 0.01% of these charge carriers flow beyond the pixel.

As shown in an alternative detailed cross-sectional view of the radiation detector 100 in FIG. 4C, according to an embodiment, the radiation absorption layer 110 may include a resistor of a semiconductor material such as, silicon, germanium, GaAs, CdTe, CdZnTe, or a combination thereof, but does not include a diode. The semiconductor may have a high mass attenuation coefficient for the radiation energy of interest.

When a radiation particle hits the radiation absorption layer 110 including a resistor but not diodes, it may be absorbed and generate one or more charge carriers by a number of mechanisms. A radiation particle may generate 10 to 100000 charge carriers. The charge carriers may drift to the electrical contacts 119A and 119B under an electric field. The field may be an external electric field. The electrical contact 119B includes discrete portions. In an embodiment, the charge carriers may drift in directions such that the charge carriers generated by a single radiation particle are not substantially shared by two different discrete portions of the electrical contact 119B (“not substantially shared” here means less than 2%, less than 0.5%, less than 0.1%, or less than 0.01% of these charge carriers flow to a different one of the discrete portions than the rest of the charge carriers). Charge carriers generated by a radiation particle incident around the footprint of one of these discrete portions of the electrical contact 119B are not substantially shared with another of these discrete portions of the electrical contact 119B. A pixel 150 associated with a discrete portion of the electrical contact 119B may be an area around the discrete portion in which substantially all (more than 98%, more than 99.5%, more than 99.9% or more than 99.99% of) charge carriers generated by a radiation particle incident therein flow to the discrete portion of the electrical contact 119B. Namely, less than 2%, less than 0.5%, less than 0.1%, or less than 0.01% of these charge carriers flow beyond the pixel associated with the one discrete portion of the electrical contact 119B.

The electronics layer 120 may include an electronic system 121 suitable for processing or interpreting signals generated by radiation particles incident on the radiation absorption layer 110. The electronic system 121 may include an analog circuitry such as a filter network, amplifiers, integrators, and comparators, or a digital circuitry such as a microprocessor, and memory. The electronic system 121 may include components shared by the pixels or components dedicated to a single pixel. For example, the electronic system 121 may include an amplifier dedicated to each pixel and a microprocessor shared among all the pixels. The electronic system 121 may be electrically connected to the pixels by vias 131. Space among the vias may be filled with a filler material 130, which may increase the mechanical stability of the connection of the electronics layer 120 to the radiation absorption layer 110. Other bonding techniques are possible to connect the electronic system 121 to the pixels without using vias.

FIG. 5A and FIG. 5B each show a component diagram of the electronic system 121, according to an embodiment. The electronic system 121 may include a first voltage comparator 301, a second voltage comparator 302, a plurality of counters 320 (including counters 320A, 320B, 320C, 320D . . . ), a switch 305, an ADC 306 and a controller 310.

The first voltage comparator 301 is configured to compare the voltage of a discrete portion of the electric contact 119B to a first threshold. The first voltage comparator 301 may be configured to monitor the voltage directly, or calculate the voltage by integrating an electric current flowing through the diode or electrical contact over a period of time. The first voltage comparator 301 may be controllably activated or deactivated by the controller 310. The first voltage comparator 301 may be a continuous comparator. Namely, the first voltage comparator 301 may be configured to be activated continuously, and monitor the voltage continuously. The first voltage comparator 301 configured as a continuous comparator reduces the chance that the system 121 misses signals generated by an incident radiation particle. The first voltage comparator 301 configured as a continuous comparator is especially suitable when the incident radiation intensity is relatively high. The first voltage comparator 301 may be a clocked comparator, which has the benefit of lower power consumption. The first voltage comparator 301 configured as a clocked comparator may cause the system 121 to miss signals generated by some incident radiation particles. When the incident radiation intensity is low, the chance of missing an incident radiation particle is low because the time interval between two successive photons is relatively long. Therefore, the first voltage comparator 301 configured as a clocked comparator is especially suitable when the incident radiation intensity is relatively low. The first threshold may be 1-5%, 5-10%, 10%-20%, 20-30%, 30-40% or 40-50% of the maximum voltage one incident radiation particle may generate on the electric contact 119B. The maximum voltage may depend on the energy of the incident radiation particle (i.e., the wavelength of the incident radiation), the material of the radiation absorption layer 110, and other factors. For example, the first threshold may be 50 mV, 100 mV, 150 mV, or 200 mV.

The second voltage comparator 302 is configured to compare the voltage to a second threshold. The second voltage comparator 302 may be configured to monitor the voltage directly, or calculate the voltage by integrating an electric current flowing through the diode or the electrical contact over a period of time. The second voltage comparator 302 may be a continuous comparator. The second voltage comparator 302 may be controllably activate or deactivated by the controller 310. When the second voltage comparator 302 is deactivated, the power consumption of the second voltage comparator 302 may be less than 1%, less than 5%, less than 10% or less than 20% of the power consumption when the second voltage comparator 302 is activated. The absolute value of the second threshold is greater than the absolute value of the first threshold. As used herein, the term “absolute value” or “modulus” |x| of a real number x is the non-negative value of x without regard to its sign. Namely,

${x} = \left\{ {\begin{matrix} {x,} & {{{if}\mspace{14mu} x} \geq 0} \\ {{- x},} & {{{if}\mspace{14mu} x} \leq 0} \end{matrix}.} \right.$ The second threshold may be 200%-300% of the first threshold. For example, the second threshold may be 100 mV, 150 mV, 200 mV, 250 mV or 300 mV. The second voltage comparator 302 and the first voltage comparator 310 may be the same component. Namely, the system 121 may have one voltage comparator that can compare a voltage with two different thresholds at different times.

The first voltage comparator 301 or the second voltage comparator 302 may include one or more op-amps or any other suitable circuitry. The first voltage comparator 301 or the second voltage comparator 302 may have a high speed to allow the system 121 to operate under a high flux of incident radiation. However, having a high speed is often at the cost of power consumption.

The counters 320 may be a software component (e.g., numbers stored in a computer memory) or a hardware component (e.g., 4017 IC and 7490 IC). Each counter 320 is associated with a bin for an energy range. For example, counter 320A may be associated with a bin for 70-71 KeV, counter 320B may be associated with a bin for 71-72 KeV, counter 320C may be associated with a bin for 72-73 KeV, counter 320D may be associated with a bin for 73-74 KeV. When the energy of an incident radiation particle is determined by the ADC 306 to be in the bin a counter 320 is associated with, the number registered in the counter 320 is increased by one.

The controller 310 may be a hardware component such as a microcontroller and a microprocessor. The controller 310 is configured to start a time delay from a time at which the first voltage comparator 301 determines that the absolute value of the voltage equals or exceeds the absolute value of the first threshold (e.g., the absolute value of the voltage increases from below the absolute value of the first threshold to a value equal to or above the absolute value of the first threshold). The absolute value is used here because the voltage may be negative or positive, depending on whether the voltage of the cathode or the anode of the diode or which electrical contact is used. The controller 310 may be configured to keep deactivated the second voltage comparator 302, the counter 320 and any other circuits the operation of the first voltage comparator 301 does not require, before the time at which the first voltage comparator 301 determines that the absolute value of the voltage equals or exceeds the absolute value of the first threshold. The time delay may expire after the voltage becomes stable, i.e., the rate of change of the voltage is substantially zero. The phase “the rate of change is substantially zero” means that temporal change is less than 0.1%/ns. The phase “the rate of change is substantially non-zero” means that temporal change of the voltage is at least 0.1%/ns.

The controller 310 may be configured to activate the second voltage comparator during (including the beginning and the expiration) the time delay. In an embodiment, the controller 310 is configured to activate the second voltage comparator at the beginning of the time delay. The term “activate” means causing the component to enter an operational state (e.g., by sending a signal such as a voltage pulse or a logic level, by providing power, etc.). The term “deactivate” means causing the component to enter a non-operational state (e.g., by sending a signal such as a voltage pulse or a logic level, by cut off power, etc.). The operational state may have higher power consumption (e.g., 10 times higher, 100 times higher, 1000 times higher) than the non-operational state. The controller 310 itself may be deactivated until the output of the first voltage comparator 301 activates the controller 310 when the absolute value of the voltage equals or exceeds the absolute value of the first threshold.

The controller 310 may be configured to cause the number registered by one of the counters 320 to increase by one, if, during the time delay, the second voltage comparator 302 determines that the absolute value of the voltage equals or exceeds the absolute value of the second threshold, and the energy of the radiation particle falls in the bin associated with the counter 320.

The controller 310 may be configured to cause the ADC 306 to digitize the voltage upon expiration of the time delay and determine based on the voltage which bin the energy of the radiation particle falls in.

The controller 310 may be configured to connect the electric contact 119B to an electrical ground, so as to reset the voltage and discharge any charge carriers accumulated on the electric contact 119B. In an embodiment, the electric contact 119B is connected to an electrical ground after the expiration of the time delay. In an embodiment, the electric contact 119B is connected to an electrical ground for a finite reset time period. The controller 310 may connect the electric contact 119B to the electrical ground by controlling the switch 305. The switch may be a transistor such as a field-effect transistor (FET).

In an embodiment, the system 121 has no analog filter network (e.g., a RC network). In an embodiment, the system 121 has no analog circuitry.

The ADC 306 may feed the voltage it measures to the controller 310 as an analog or digital signal. The ADC may be a successive-approximation-register (SAR) ADC (also called successive approximation ADC). An SAR ADC digitizes an analog signal via a binary search through all possible quantization levels before finally converging upon a digital output for the analog signal. An SAR ADC may have four main subcircuits: a sample and hold circuit to acquire the input voltage (V_(in)), an internal digital-analog converter (DAC) configured to supply an analog voltage comparator with an analog voltage equal to the digital code output of the successive approximation register (SAR), the analog voltage comparator that compares V_(in) to the output of the internal DAC and outputs the result of the comparison to the SAR, the SAR configured to supply an approximate digital code of V_(in) to the internal DAC. The SAR may be initialized so that the most significant bit (MSB) is equal to a digital 1. This code is fed into the internal DAC, which then supplies the analog equivalent of this digital code (V_(ref)/2) into the comparator for comparison with V_(in). If this analog voltage exceeds V_(in) the comparator causes the SAR to reset this bit; otherwise, the bit is left a 1. Then the next bit of the SAR is set to 1 and the same test is done, continuing this binary search until every bit in the SAR has been tested. The resulting code is the digital approximation of V_(in) and is finally output by the SAR at the end of the digitization.

The system 121 may include a capacitor module 309 electrically connected to the electric contact 119B, wherein the capacitor module is configured to collect charge carriers from the electric contact 119B. The capacitor module can include a capacitor in the feedback path of an amplifier. The amplifier configured as such is called a capacitive transimpedance amplifier (CTIA). CTIA has high dynamic range by keeping the amplifier from saturating and improves the signal-to-noise ratio by limiting the bandwidth in the signal path. Charge carriers from the electrode accumulate on the capacitor over a period of time (“integration period”) (e.g., as shown in FIG. 6 , between t_(s) to t₀). After the integration period has expired, the capacitor voltage is sampled by the ADC 306 and then reset by a reset switch. The capacitor module 309 can include a capacitor directly connected to the electric contact 119B.

FIG. 6 schematically shows a temporal change of the electric current flowing through the electric contact 119B (upper curve) caused by charge carriers generated by a radiation particle incident on the pixel 150 associated with the electric contact 119B, and a corresponding temporal change of the voltage of the electric contact 119B (lower curve). The voltage may be an integral of the electric current with respect to time. At time to, the radiation particle hits the diode or the resistor, charge carriers start being generated in the pixel 150, electric current starts to flow through the electric contact 119B, and the absolute value of the voltage of the electric contact 119B starts to increase. At time t₁, the first voltage comparator 301 determines that the absolute value of the voltage equals or exceeds the absolute value of the first threshold V1, and the controller 310 starts the time delay TD1 and the controller 310 may deactivate the first voltage comparator 301 at the beginning of TD1. If the controller 310 is deactivated before t₁, the controller 310 is activated at t₁. During TD1, the controller 310 activates the second voltage comparator 302. The term “during” a time delay as used here means the beginning and the expiration (i.e., the end) and any time in between. For example, the controller 310 may activate the second voltage comparator 302 at the expiration of TD1. If during TD1, the second voltage comparator 302 determines that the absolute value of the voltage equals or exceeds the absolute value of the second threshold at time t₂, the controller 310 waits for stabilization of the voltage to stabilize. The voltage stabilizes at time t_(e), when all charge carriers generated by the radiation particle drift out of the radiation absorption layer 110. At time t_(s), the time delay TD1 expires. At or after time t_(e), the controller 310 causes the ADC 306 to digitize the voltage and determines which bin the energy of the radiation particle falls in. The controller 310 then causes the number registered by the counter 320 corresponding to the bin to increase by one. In the example of FIG. 6 , time t_(s) is after time t_(e); namely TD1 expires after all charge carriers generated by the radiation particle drift out of the radiation absorption layer 110. If time t_(e) cannot be easily measured, TD1 can be empirically chosen to allow sufficient time to collect essentially all charge carriers generated by a radiation particle but not too long to risk have another incident radiation particle. Namely, TD1 can be empirically chosen so that time t_(s) is empirically after time t_(e). Time t_(s) is not necessarily after time t_(e) because the controller 310 may disregard TD1 once V2 is reached and wait for time t_(e). The rate of change of the difference between the voltage and the contribution to the voltage by the dark current is thus substantially zero at t_(e). The controller 310 may be configured to deactivate the second voltage comparator 302 at expiration of TD1 or at t₂, or any time in between.

The voltage at time t_(e) is proportional to the amount of charge carriers generated by the radiation particle, which relates to the energy of the radiation particle. The controller 310 may be configured to determine the bin the energy of the radiation particle falls in, based on the output of the ADC 306.

After TD1 expires or digitization by the ADC 306, whichever later, the controller 310 connects the electric contact 119B to an electric ground for a reset period RST to allow charge carriers accumulated on the electric contact 119B to flow to the ground and reset the voltage. After RST, the system 121 is ready to detect another incident radiation particle. Implicitly, the rate of incident radiation particles the system 121 can handle in the example of FIG. 6 is limited by 1/(TD1+RST). If the first voltage comparator 301 has been deactivated, the controller 310 can activate it at any time before RST expires. If the controller 310 has been deactivated, it may be activated before RST expires.

Because the radiation detector 100 has several pixels 150 that may operate in parallel, the radiation detector can handle much higher rate of incident radiation particles. This is because the rate of incidence on a particular pixel 150 is 1/N of the rate of incidence on the entire array of pixels, where N is the number of pixels.

FIG. 7 shows an example flow chart for using the endoscope 101 with the radiation detector 100, according to an embodiment. In step 701, the endoscope 101 is inserted into a human body. In step 702, an intensity of radiation in a first range of energy and an intensity of radiation in a second range of energy are detected using the radiation detector 100. The radiation may be emitted from a tissue (e.g., blood vessel) inside the human body due to radiation fluorescence as a result of exposure of the tissue to radiation from outside the human body. In step 703, the existence of a chemical (e.g., Ca) in the human body is determined based on the intensities. For example, the first range of energy may encompass Kα emission peak of calcium and the second range of energy may encompass Kβ emission peak of calcium. If the intensities have the ratio characteristics to X-ray fluorescence from calcium, the existence of calcium in the vicinity of the radiation detector 100 may be determined.

FIG. 8 schematically shows a system including the endoscope 101 described above and a radiation source 1601. The radiation source 1601 may be configured to induce radiation fluorescence of a tissue in a human body 1602. The system may be used for detecting and determining concertation of calcification inside blood vessels. The radiation source 1601 may be configured to emit X-ray or gamma ray.

While various aspects and embodiments have been disclosed herein, other aspects and embodiments will be apparent to those skilled in the art. The various aspects and embodiments disclosed herein are for purposes of illustration and are not intended to be limiting, with the true scope and spirit being indicated by the following claims. 

What is claimed is:
 1. An endoscope, comprising: an insertion tube; a radiation detector configured to detect radiation particles in a first range of energy and radiation particles in a second range of energy; wherein the radiation detector is configured to determine existence of a chemical element based on intensity of the radiation particles in the first range of energy and intensity of the radiation particles in the second range of energy.
 2. The endoscope of claim 1, wherein the radiation particles are X-ray photons.
 3. The endoscope of claim 1, wherein the radiation particles are emitted by fluorescence of a tissue in a person.
 4. The endoscope of claim 3, wherein the tissue is a blood vessel.
 5. The endoscope of claim 1, wherein the radiation detector is configured to count numbers of radiation particles incident thereon whose energy fall in a plurality of bins, within a period of time.
 6. The endoscope of claim 1, wherein the chemical element is calcium.
 7. The endoscope of claim 1, wherein the radiation detector is configured to determine a spectrum of the radiation particles incident on the radiation detector.
 8. The endoscope of claim 1, wherein the radiation detector comprises a plurality of pixels, wherein each of the pixels is configured to count numbers of radiation particles incident thereon whose energy fall in a plurality of bins, within a period of time; and wherein the radiation detector is configured to determine a sum of the numbers of radiation particles for the bins of the same energy range counted by all the pixels.
 9. The endoscope of claim 8, wherein the pixels are configured to count the numbers of X-ray photons within a same period of time.
 10. The endoscope of claim 8, wherein each of the pixels comprises an analog-to-digital converter (ADC) configured to digitize an analog signal representing the energy of an incident radiation particle into a digital signal.
 11. The endoscope of claim 8, wherein the pixels are configured to operate in parallel.
 12. The endoscope of claim 10, wherein the ADC is a successive-approximation-register (SAR) ADC.
 13. The endoscope of claim 1, wherein the radiation detector comprises: a radiation absorption layer comprising an electric contact; a first voltage comparator configured to compare a voltage of the electric contact to a first threshold; a second voltage comparator configured to compare the voltage to a second threshold; a controller; a plurality of counters each associated with a bin and configured to register a number of radiation particles absorbed by the radiation absorption layer wherein the energy of the radiation particles falls in the bin; wherein the controller is configured to start a time delay from a time at which the first voltage comparator determines that an absolute value of the voltage equals or exceeds an absolute value of the first threshold; wherein the controller is configured to determine whether an energy of a radiation particle falls into the bin; wherein the controller is configured to cause the number registered by the counter associated with the bin to increase by one.
 14. The endoscope of claim 13, further comprising a capacitor module electrically connected to the electric contact, wherein the capacitor module is configured to collect charge carriers from the electric contact.
 15. The endoscope of claim 13, wherein the controller is configured to activate the second voltage comparator at a beginning or expiration of the time delay.
 16. The endoscope of claim 13, wherein the controller is configured to connect the electric contact to an electrical ground.
 17. The endoscope of claim 13, wherein a rate of change of the voltage is substantially zero at expiration of the time delay.
 18. The endoscope of claim 13, wherein the radiation absorption layer comprises a diode.
 19. The endoscope of claim 13, wherein the radiation absorption layer comprises silicon, germanium, GaAs, CdTe, CdZnTe, or a combination thereof.
 20. The endoscope of claim 1, wherein the radiation detector does not comprise a scintillator.
 21. A system comprising the endoscope of claim 1, and a radiation source.
 22. A method comprising: inserting an endoscope of claim 1 into a human body; detecting, using the radiation detector, an intensity of radiation in a first range of energy and an intensity of radiation in a second range of energy; determining existence of a chemical element in the human body based on the intensity of radiation in the first range of energy and the intensity of radiation in the second range of energy. 